The field of the invention is nuclear magnetic resonance imaging methods and systems. More particularly, the invention relates to fast NMR pulse sequences with enhanced susceptibility weighting in the reconstructed image.
Any nucleus that possesses a magnetic moment attempts to align itself with the direction of the magnetic field in which it is located. In doing so, however, the nucleus precesses around this direction at a characteristic angular frequency (Larmor frequency) which is dependent on the strength of the magnetic field and on the properties of the specific nuclear species (the magnetogyric constant .gamma. of the nucleus). Nuclei that exhibit this phenomena are referred to herein as "spins".
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B.sub.0), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. A net magnetic moment M.sub.z is produced in the direction of the polarizing field, but the randomly oriented magnetic components in the perpendicular, or transverse, plane (x-y plane) cancel one another. If, however, the substance, or tissue, is subjected to a magnetic field (excitation field B.sub.1) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, M.sub.z, may be rotated, or "tipped", into the x-y plane to produce a net transverse magnetic moment M.sub.t, which is rotating, or spinning, in the x-y plane at the Larmor frequency. The practical value of this phenomenon resides in the signal that is emitted by the excited spins after the excitation signal B.sub.1 is terminated. There is a wide variety of measurement sequences in which this nuclear magnetic resonance ("NMR") phenomena is exploited.
In simple systems the excited spins induce an oscillating sine wave signal in a receiving coil. The frequency of this NMR signal is the Larmor frequency, and its initial amplitude, A.sub.0, is determined by the magnitude of the transverse magnetic moment M.sub.xy. The amplitude, A, of the emission signal decays in an exponential fashion with time, t: EQU A=A.sub.0 e.sup.-t/T*.sbsp.2
The decay constant 1/T*.sub.2 is inversely proportional to the exponential rate at which the aligned precession of the spins dephase after removal of the excitation signal B.sub.1. This dephasing is caused by local magnetic field inhomogeneities produced in part by the differences in susceptibility between the spins. The dephasing caused by the spin system itself is referred to as the "spin-spin relaxation constant" T.sub.2, and as will be described below, this characteristic is used in medical imaging to contrast tissues containing spins that exhibit different spin-spin relaxation times. NMR images that rely on this susceptibility phenomenon to contrast different tissues are referred to as "T*.sub.2 weighted images."
When utilizing NMR to produce images, a technique is employed to obtain NMR signals from specific locations in the subject. Typically, the region which is to be imaged (region of interest) is scanned by a sequence of NMR measurement cycles which vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques. To perform such a scan, it is, of course, necessary to elicit NMR signals from specific locations in the subject. This is accomplished by employing magnetic fields (G.sub.x, G.sub.y, and G.sub.z) which have the same direction as the polarizing field B.sub.0, but which have a gradient along the respective x, y and z axes. By controlling the strength of these gradients during each NMR cycle, the spatial distribution of spin excitation can be controlled and the location of the resulting NMR signals can be identified.
Most NMR scans currently used to produce medical images require many minutes to acquire the necessary data. The reduction of this scan time is an important consideration, since reduced scan time increases patient throughput, improves patient comfort, and improves image quality by reducing motion artifacts. In addition, if body functions such as brain activity are to be monitored with a series of images, it is imperative that each image be acquired in seconds, rather than minutes.
There is a class of fast pulse sequences that have a very short repetition time (TR) and result in complete scans which can be conducted in seconds rather than minutes. Whereas the more conventional pulse sequences have repetition times TR which are much greater than the spin-spin relaxation constant T.sub.2 so that the transverse magnetization has time to relax between the phase coherent excitation pulses in successive sequences, the fast pulse sequences have a repetition time TR which is less than T.sub.2 and which drives the transverse magnetization into a steady-state of equilibrium. Such techniques are referred to as steady-state free precession (SSFP) techniques and they are characterized by a cyclic pattern of transverse magnetization in which the resulting NMR signal refocuses at each RF excitation pulse to produce an echo signal. This echo signal includes a first part S+ that is produced after each RF excitation pulse and a second part S- which forms just prior to the RF excitation pulse.
One well known SSFP pulse sequence is called gradient refocused acquired steady-state (GRASS). It utilizes a readout gradient G.sub.x to shift the peak in the S+ signal that is produced after each RF excitation pulse toward the center of the pulse sequence. In two-dimensional imaging, a slice selection gradient pulse is produced by the gradient G.sub.z and is immediately refocused in the well-known manner. A phase encoding gradient pulse G.sub.y is produced shortly thereafter to position encode the acquired NMR data. By using this SSFP pulse sequence, data for a complete image can be acquired in less than two seconds and a series of such images can be acquired over a period of time which enable such processes as brain functions to be dynamically monitored. Each acquired image captures the state of the monitored function over a relatively short time interval, and a "time resolution" of less than two seconds per image can be achieved.
As indicated above, a very useful image contrast mechanism in NMR medical imaging is T*.sub.2 contrast. As described by Peter A. Bandettini et al in "Time Course EPI of Human Brain Function During Task Activation," Magnetic Resonance in Medicine, 25, 390-397 (1992), the paramagnetic characteristics of blood change as a function of its oxygenation and this is reflected as a change in its T*.sub.2 constant. Thus, a series of NMR images may show the identical physical structures, but the brightness of the perfused tissue will differ as a function of the degree of oxygenation of the blood in the capillary bed. The amount of image contrast depends on the extent to which the T*.sub.2 changes. In turn, the amount of change in T*.sub.2 can be enhanced by using paramagnetic contrast agents such as Gadolinium DPTA, or by increasing the T*.sub.2 sensitivity of the pulse sequence by lengthening the echo time (TE).
While SSFP pulse sequences provide the desired time resolution for dynamic monitoring, they are not T*.sub.2 contrast sensitive because of their very short echo times (TE). Consequently, other pulse sequences, such as echo planar sequences (EPI), have been employed for dynamic studies because they provide good time resolution and good T*.sub.2 contrast. Unfortunately, EPI pulse sequences often require hardware enhancements to standard NMR systems and their use is, therefore, limited.